Micro blood vessels and tissue ducts

ABSTRACT

A fiber includes one or more layers of polymer surrounding a central lumen, and living animal cells disposed within the lumen and/or within at least one of the one or more layers, wherein the fiber has an outer diameter of between 5 and 8000 microns and wherein each individual layer of polymer has a thickness of between 0.1 and 250 microns. Also disclosed are model tissues including such fibers, and method of making such fibers. The fibers can serve as synthetic blood vessels, ducts, or nerves.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a Division of U.S. patent application Ser. No.13/784,216 filed on Mar. 4, 2013 which is a Continuation-In-Part of U.S.patent application Ser. No. 13/081,688 filed on Apr. 7, 2011 (now U.S.Pat. No. 8,398,935), which in turn is a Continuation-In-Part of U.S.patent application Ser. No. 11/423,225 filed on Jun. 9, 2006 (now U.S.Pat. No. 8,361,413), which claims the benefit of U.S. Provisional PatentApplication Ser. No. 60/690,057 filed on Jun. 9, 2005. This applicationis also related to U.S. patent application Ser. Nos. 12/987,251 and13/728,651, filed on Jan. 10, 2011 and Dec. 27, 2012, respectively. Eachof these applications is incorporated herein by reference.

BACKGROUND

Sheath flow is a widely used technique for a variety of applications,including but not limited to particle counting, flow cytometry,waveguiding, and fluid control. Sheath flow involves surrounding acentral flow stream (the core) with a surrounding stream (the sheath).In particle counting and flow cytometry applications, the sheathprevents particles in the core from coming into contact with the wallsof the channel, thus preventing adhesion and clogging. The sheath alsoserves to focus the particles or molecules into the center of thechannel, allowing for easy counting or measurement through optical orother means. Sheath flow is normally laminar flow that substantiallyavoids mixing between the core stream and the sheath stream. Sheath flowcan also be used with fluids of different refractive index to create awaveguide in the core or sheath stream in order to measure transfer ofanalytes from one stream to the other or to control the rate ofinteraction between molecules in one or both streams for carefullycontrolled chemistry or analysis.

Previous designs have created sheath flow through an annulararrangement. A small nozzle was positioned inside a larger tube. Thecore solution was pumped through the nozzle and the sheath solution waspumped through the larger tube. This configuration required carefulalignment of the two tubes and did not easily lend itself tominiaturization. Since the diameter of the nozzle was fixed, therelative sizes of the core stream and sheath solution were relativelyconstant within a set range.

Other devices provide sheath flow on a chip, but the flow typicallyoperates only in two dimensions. The core stream in these devices isbordered on either side by the sheath streams, however the core is notsheathed top and bottom. The complexity of the support plumbing forthese devices is increased, as the number of flow streams is increasedfrom two to three as compared to the annular arrangement design. It ispossible to sheath the stream on the top and bottom of the core streamin these systems by adding two additional inlet ports in the top andbottom of the channel. However, this greatly increases the manufacturingcomplexity of the device. Micromachining technologies are inherentlytwo-dimensional. Three-dimensional channel paths can be created bystacking several two dimensional designs on top of one another, but thisadds to the complexity and difficulty of the manufacturing process.Creating a fully sheathed flow in this way could require at leastseveral individual levels, which must be independently produced and thencarefully aligned. In addition, use of the device could require multiplepumps to provide solutions to all the inlets.

Tissue Engineering

Studying the growth and differentiation of cells in culture in thepresence of various nutrients and growth factors has providedphysiologically relevant information about how the corresponding cellsfunction in vivo. Classic tissue engineering mixes cell types withdifferent growth factors and provides minimal control over morphologyand microanatomy [Langer, R., Vacanti, J. P. 1993 “Tissue Engineering.”Science. 260(5110):920-6; Levenberg, S., Rouwkema, J., Macdonald, M.,Garfein, E. S., Kohane, D. S., Darland, D. C., et al. 2005 “Engineeringvascularized skeletal muscle tissue.” Nat Biotechnol. 23(7):879-84.]. Itis becoming increasingly clear that cell differentiation and functionare impacted by fluid flow, proximity of other cell types, and substrategeometry. The realization of the importance of such factors hasmotivated the microfluidics and tissue engineering communities to create“tissue-on-chip” or “organ-on-chip” model systems that can introducemethods for controlling such variables and provide more complex in vitrosystems for the study of normal differentiation and pathogenesis or drugmetabolism and transport (Wong, K. H. K., Chan, J. M., Kamm, R. D.,Tien, J. 2012 Microfluidic Models of Vascular Functions. Annu Rev BiomedEng. 14:205-230; Van Der Meer, A. D., Van Den Berg, A. 2012“Organs-on-chips: breaking the in vitro impasse.” Integr Biol-UK4(5):461-470; Shuler, M. L. 2012 “Modeling Life.” Ann. Biomed. Eng.40(7):1399-1407).

In general, these tissue-on-chip models are configured with one of twotypes of architectures. For decades, investigators have been configuringsubstrates to have a specific geometry that will impact celldifferentiation, most notably defining surface topography or creatingchannels to direct cell growth. In both cases, the cells are introducedafter the substrate is configured, and cells adhere to defined portionsof the substrate. Surface patterning of endothelial or progenitor cellshas been used to engineer blood vessels [Niklason, L. E., Gao, J.,Abbott, W. M., Hirschi, K. K., Houser, S., Marini, R., et al. 1999“Functional arteries grown in vitro.” Science. 284(5413):489-93;Kaushal, S., Amiel, G. E., Guleserian, K. J., Shapira, O. M., Perry, T.,Sutherland, F. W., et al. 2001 “Functional small-diameter neovesselscreated using endothelial progenitor cells expanded ex vivo.” Nat Med.7(9):1035-40.]. Chip-based approaches to generating engineered bloodvessels pattern layers of cells onto tubular or rectangularmicrochannels to emulate blood vessel geometry; however, these methodsdo not produce a free-standing engineered blood vessel. Themicrochannel-attached engineered blood vessel does not allow forsuperfusion along the blood vessel wall or branching from the mainvessel. [Nichol, J. W., Koshy, S. T., Bae, H., Hwang, C. M., Yamanlar,S., Khademhosseini, A. 2010 “Cell-laden microengineered gelatinmethacrylate hydrogels.” Biomaterials. 31(21):5536-44. Chau, L. T.,Rolfe, B. E., Cooper-White, J. J. 2011 “A microdevice for the creationof patent, three-dimensional endothelial cell-based microcirculatorynetworks.” Biomicrofluidics. 5(3); Liu, Y. X., Markov, D. A., Wikswo, J.P., Mccawley, L. J. 2011 “Microfabricated scaffold-guided endothelialmorphogenesis in three-dimensional culture.” Biomed Microdevices.13(5):837-46.] More recently, porous membranes have been suspendedacross a microfluidic well or channel and cells grown on one or bothsurfaces of the porous membrane. The cells generally form monolayerswhile air or liquids are flowed above and/or below the membrane (forexample, U.S. Patent Application Publication No. 2011/0250585A1 and Huh,D, Ingber et al., “Reconstituting organ-level lung functions on a chip,”Science, 2010, 328, 1662.). Both of these approaches limit thetissue-on-chip construct to planar configurations for the resulting cellorganizations. The depth of the cell layers in planar tissue models isconstrained by the need to transport nutrients and growth factors fromthe fluid in the microfluidic channel, preventing the formation of thicktissues.

In nature, the transport of nutrients, growth factors, protectivemolecules, and waste are provided by the vasculature and other types ofducts. Cells including both those that provide immunity and oxygenationand those that cause infection, autoimmune disease and cancer are alsotransported through the vasculature. The need for tubular structures toincorporate into tissue-on-chip models is recognized (e.g. Wong et al.,ibid.). Ideally, such vasculature would be round as in nature, flexibleto accommodate complex organ geometries, and composed of a biocompatibleor biodegradable material that can be remodeled by the incorporatedcells (not polydimethylsiloxane). The roadblock to vascularized tissuemodels is twofold: (1) accurate fabrication of engineered blood vesselsand (2) integration of engineered blood vessels into on-chip models.

In the present state of the art, elongated blood vessel structuresrequire either cells grown in channels [Franco, C., Gerhardt, H. 2012“Tissue Engineering: Blood vessels on a chip.” Nature. 488(7412):465-6],cells seeded onto preformed scaffolds, or cultures using blood vesselsexcised from animals [Quint, C., Kondo, Y., Manson, R. J., Lawson, J.H., Dardik, A., Niklason, L. E. 2011 “Decellularized tissue-engineeredblood vessel as an arterial conduit.” Proc. Nat'l Acad. Sci. USA.108(22):9214-9]. The first methods do not provide stand-alone bloodvessels and culturing excised vessels is not a viable source forproducing cost and time-effective human tissue models. Production ofcollagen-based blood vessel scaffolds for subsequent cell incorporationis already being used clinically, where it is desirable for the patientto provide his own cells. However, these scaffolds are large and replaceveins or arteries rather than capillaries. The scaffold structures aregenerally millimeters in diameter, and nutrient availability in vitromight limit cells to growth near the surface of the scaffold, makingthem less appropriate for tissue-on-chip models. Development of methodsfor producing blood vessels to operate and supply tissue models wouldprovide a much more accurate model for blood delivery to tissue andprovide for growth of on-chip tissues in three dimensions.

The range of applications for the nano- and microfabrication ofcell-laden and/or coated tubular constructs are not limited tovasculature. Directed nerve growth has been shown whereby neurons areseeded into scaffolds composed of poly L-lactic acid nanofibers. [F.Yang, R. Murugan, S. Ramakrishna, X. Wang, Y.-X. Ma, S. Wang,“Fabrication of nano-structured porous PLLA scaffold intended for nervetissue engineering,” Biomaterials 25 (2004) 1891-1900]. The scaffoldsinitiated axonal guidance, which was postulated as a first step inbridging the gaps between the proximal and distal nerves that do notclose during healing. An inkjet printing station for neuroregenerativetissue engineering was presented by Silva, D. S., D. B. Wallace, et al.(2007) [IEEE Dallas Engineering in Medicine and Biology Workshop:71-73]. The resulting tubes were seeded with neurons and served asscaffolds that resulted in significant outgrowth.

Ductal tissues play essential roles in human physiology by providingconduits that facilitate the transfer of fluids such as seminal fluid,bile, and milk. These conduits serve as the interfaces between exocrineglands and distal regions both internal and external. Ducts differ fromblood vessels in that they are often not simply passive conduits, butactively participate in the secretions that facilitate the expelling ofwaste or transfer of reproductive materials. As an example, a multilayermicrofluidic device that was used to culture and analyze renal tubulecells [Jang, K.-J., and Suh, K.-Y. “A multi-layer microfluidic devicefor efficient culture and analysis of tubular cells.” Lab on a Chip(2010) 10, 36-42]. A fibronectin-coated polyester membrane was insertedbetween the layers of a polydimethylsiloxane (PDMS) microfluidicchannel, and primary kidney cells were introduced on one side. Duringculture, the membrane was subjected to continuous shear stress of 1dyn/cm² for 5 h. The cells formed a layer on the polyester membrane anddeveloped markers typical of renal tubule cells.

In view of the above, a need exists for the engineering of bloodvessels, tissue ducts, and the like having physiologically-appropriateshapes, dimensions, and properties. Most prior methods for makingpolymer fibers involve conditions that are not compatible with livingcells, in that they typically involve elevated temperature, high sheer,organic solvents, or combinations of these. Techniques described hereinprovide a biocompatible method for making shaped polymer fibers usinghydrodynamic focusing.

BRIEF SUMMARY

A fiber comprises one or more layers of polymer surrounding a centrallumen, and living mammalian cells disposed within the lumen and/orwithin at least one of the one or more layers, wherein the fiber has anouter diameter of between 5 and 8000 microns and wherein each individuallayer of polymer has a thickness of between 0.1 and 250 microns. Thefibers is in a condition of having been generated via sheath flow. Thelumen is optionally hollow or filled with a polymer.

Another embodiment is a model tissue comprising an inlet port and anoutlet port, and at least one fiber comprising: one or more layers ofpolymer surrounding a central lumen, and living mammalian cells disposedwithin the lumen and/or within at least one of the one or more layers,wherein the fiber has an outer diameter of between 5 and 8000 micronsand wherein each individual layer of polymer has a thickness of between0.1 and 250 microns, and wherein the fiber is in a condition of havingbeen generated via sheath flow.

A further embodiment is a method of generating a fiber by creating asheath flow comprising a core stream surrounded by one or more sheathstreams, wherein at least one of the core or sheath streams comprises apolymerizable material and wherein at least one of the core or sheathstreams comprises living mammalian cells; and polymerizing thepolymerizable material to form a fiber wherein the fiber comprises: oneor more layers of polymer derived from the one or more sheath streamssurrounding a central lumen derived from the core stream, and the livingmammalian cells disposed within the lumen and/or within at least one ofthe one or more layers, wherein the fiber has an outer diameter ofbetween 5 and 8000 microns and wherein each individual layer of polymerhas a thickness of between 0.1 and 250 microns. In the sheath flow, thepolymerizable material and the cells may be located in the same ordifferent streams.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a view of one example of a sheath flow device;

FIG. 2 is a view of one example of a sheath flow device;

FIGS. 3A-3F show a series of representative cross sections of sheathedflow;

FIG. 4 is a view of one example of a sheath flow device;

FIG. 5 is a view of one example of a sheath flow device;

FIG. 6 is a representative cross section of sheathed flow;

FIG. 7 is a representative cross section of sheathed flow;

FIG. 8 is a representative cross section of sheathed flow;

FIG. 9 is a representative cross section of sheathed flow;

FIG. 10 is a representative cross section of unsheathed flow;

FIGS. 11A-11C show a series of representative cross sections of sheathedflow;

FIG. 12 is a liquid waveguide device;

FIG. 13 is a representation of waveguided light though a liquidwaveguide;

FIG. 14 is a near field microscope;

FIG. 15 is a flow cytometer device;

FIGS. 16A through 16E shows the results of tests using the flowcytometer device;

FIG. 17 is a view of one example of a sheath flow device;

FIG. 18 is a tube within a tube made by the sheath flow device.

FIG. 19a is view of one example of a sheath flow device showing fluidtransporting structure across the top surface and a second fluidtransporting structure across the bottom surface.

FIG. 19b is view of one example of a sheath flow device showing fluidtransporting structure across the top surface and a second fluidtransporting structure across the bottom surface.

FIG. 20 shows how parallel core streams may be used for simultaneousshaping of two core streams. Polymerizable material is input from twoouter (left) sides of the main channel and the sheath fluid is inputinto the center of the main channel. The grooves channel thepolymerizable material and sheath solution so that the prepolymer flowsremain separate, but are completely surrounded with sheath fluid.Following polymerization, the result will be two parallel shaped fibersexiting from the same channel.

FIGS. 21A and 21B show how a split core may be obtained by controllingthe relative flow rates of the core and sheath. A single stream ofpolymerizable material can be split into multiple streams using theappropriate combination of wall structures and relative flow rates. Thesimulation of FIG. 21A shows a channel with 7 chevrons in the top andbottom and prepolymer and sheath flow through the channel at flow-rateratios of (a) 1:1, (b) 50:1, (c) 500:1, (d) 1000:1, and (e) 2500:1. FIG.21B shows a fiber run at the 2500:1 ratio which split into two filamentsthat hardened as independent, but parallel filaments.

FIGS. 22A and 22B shows a simulation of how flow through a 5-chevrondevice can be used to split a single core into two parallel streams.FIG. 22A shows the model results on the z and y axes while FIG. 22Bshows a perspective view generated by the model.

FIG. 23 shows how structures in a channel can be used to split a singlecore stream (black, input to center of a channel) into multiple streamsfor production of multiple parallel fibers.

FIG. 24 shows a fiber with variable dimensions. The fiber was cast fromacrylate in a grooved fluidic channel with variable pump pressure on theinlets to alter the flow-rate ratio of core and sheath streams.

FIG. 25 shows examples of organizing shaped fibers.

FIG. 26A shows example of multi-component fiber cross sections. FIG. 26Bshows a schematic cross-section of a fiber with a gradientcross-section.

FIG. 27 shows a curly fiber.

FIG. 28 shows a schematic cross-section of a fiber with high surfacearea.

FIG. 29 shows a fiber exhibiting regular viscous buckling.

FIG. 30 schematically illustrates the cover and substrate layers of asheath flow device used to manufacture hollow (or multi-layer/component)polymer fibers.

FIGS. 31A and 31B show hollow PEG fibers manufactured using the deviceshown in FIG. 30.

FIG. 32 illustrates the cover and substrate layers of a sheath flowdevice used to manufacture hollow or solid multi-layer polymer fibers.This device operates similarly to the device presented in FIG. 30,except sheath is introduced along with a corresponding set of chevronsthat serve to add a third layer. The cross-sectional areas of flows areproduced from two sets of chevron shaping features, three sets ofchevron shaping features, and two sets of striped shaping features.

FIG. 33 shows a hydrodynamic shaping device with a representativecomputational model of the layered fluid flow. Discrete regions of flowwill correlate to micro-blood vessel layers. Each layer requires asegment of the microfluidic channel to incorporate an additionalhydrodynamic shaping device.

FIG. 34 shows computational fluid dynamic simulations of hollow tubesthat can be fabricated using hydrodynamic focusing with fluidtransporting structures in the top and bottom of microchannels. Use ofeither stripes or chevrons will produce multilayered fibers. Additionalsets of chevrons produce additional concentric layers.

FIG. 35 illustrates a model issue microchip vascularized with themicro-blood vessels that can be utilized for both mechanical andphysiological analyses. The integrated tissue reservoir incorporates a3-dimensional hydrogel matrix to support vasculogenesis and3-dimensional culturing of osteogenic cells.

DETAILED DESCRIPTION Definitions

Before describing the present invention in detail, it is to beunderstood that the terminology used in the specification is for thepurpose of describing particular embodiments, and is not necessarilyintended to be limiting. Although many methods, structures and materialssimilar, modified, or equivalent to those described herein can be usedin the practice of the present invention without undue experimentation,the preferred methods, structures and materials are described herein. Indescribing and claiming the present invention, the following terminologywill be used in accordance with the definitions set out below.

As used in this specification and the appended claims, the singularforms “a,” “an,” and “the” do not preclude plural referents, unless thecontent clearly dictates otherwise.

As used herein, the term “and/or” includes any and all combinations ofone or more of the associated listed items.

As used herein, a “core” refers to a fluid flow that is concentricallysurrounded by another fluid flow, termed the “sheath.” Together the coreand sheath flow are referred to as a “sheathed flow.” A core mayoptionally include within it an interior core, so that the interior coresurrounded by an exterior portion of the core serving as a sheath.Optionally, the interior core may in turn serve as a sheath to a deeperinterior core, and so on. The cores may have differing compositions. Asused herein, the term “simple core” refers to a core lacking an interiorcore.

As used herein, the term “cross-section” refers to cross-sectionalshape, area, and/or dimension(s).

Description

In the present device and method, one or more core streams and one ormore sheath streams are introduced into a single channel. One or morefluid transporting structures located at the top and bottom of thechannel direct the sheath fluid around the core stream, separating thecore stream from the walls of the channel. Once the position of the corestream is established in the interior of the channel, it remains in thatposition due to laminar flow.

FIG. 1 shows a top view of one example of a sheath flow device. A sheathstream inlet 10, and a core stream inlet 12, allow a sheath stream and acore stream to be introduced into a channel 14. One design provides fora at a ‘T’ intersection at the proximal end 16 of the channel 14. Thesheath stream and the core stream flow down the channel side-by-sidetowards the distal end of the channel 18 where an outlet 20 is present.At least one fluid transporting structure 22 such as a groove or a ridgeis located in the channel 14 between the inlets 10, 12 and the outlet20. The fluid transporting structure 22 transports the sheath streamacross the top and bottom of the channel 14 to completely surround thecore stream. The fluid transporting structure 22 crosses the channel 14at an angle 30.

The device can be readily fabricated using a variety of techniques,including molding, milling, laser ablation, soft lithography techniquesand other fabrication techniques known to those skilled in the art. Anymaterial that can be machined or molded into the appropriate shapes canbe used. The current techniques used in the mass production ofmicrofluidic components can be easily adapted to the production of thissheath flow design.

The exact shape of the channel is not critical. For example, FIG. 2shows a channel 14 with a constriction at the location of the grooves22. The constricted device showed similar behavior to devices withoutthe constriction. The size of the channel can be varied within a broadrange of size scales. The size of the channel is limited at the lowerend by diffusion. When the width or diameter of the channel reaches thediffusional distance of the molecules or particles of interest, anyattempts to confine them to a specific region of the channel will bethwarted.

The upper limit for the channel width is set by the Reynolds number ofthe system. The device shown in FIG. 1 has been shown to function atReynolds numbers up to and including 200. This means that the device canbe fabricated into larger sizes using slower velocities or higherviscosity fluids. Sheath flow devices have been fabricated for use withhigh viscosity fluids that are 3 mm in width that have Reynolds numbersof 0.0008, so the actual channel diameter can be significantly widerthan that with the use of appropriate fluids. The device will operate atReynolds numbers up to those at which turbulence is initiated.

The channel has at least two inlets at or near its proximal end. Theinlets are used to introduce a sheath stream and a core stream into thechannel. The size and exact location of the inlets are can be varied,provided that the fluid transporting structure in the channel is locateddownstream from the inlets.

The at least one fluid transporting structure is typically a groove or aridge located inside the channel. The structure transports the sheathstream laterally across the channel and around the core stream,separating the core stream from the walls of the channel. Once theposition of the core stream is established in the interior of thechannel, it remains in that position due to laminar flow. The angle ofthe fluid transporting structure across the channel is not necessarilycritical to the design, however it has been found to be important inapplications involving shaping of the core. FIG. 1 shows a device havinga fluid transporting structure 22 that has an angle 30 that is about 45°relative to the channel; however, other oblique angles will work aswell.

The number and depth of the fluid transporting structures are designparameters that also can be adjusted to suit particular applications. Asingle structure located on the top and bottom of the channel willprovide for a full sheath around the core stream. The grooves do nothave to be precisely aligned along the flow axis in order for the deviceand method to operate. However, their lateral alignment may beimportant. Increasing the number of fluid transporting structuresprovides control over the lateral position of the core within thechannel. Increasing the size of the fluid transporting structurescorrelates with a more effective transport of the sheath stream acrossthe channel. Preferably, the fluid transporting structures penetrate thewall of the channel on the downstream end. FIG. 1 shows the fluidtransporting structures 22 penetrating the wall of the channel 14 on thedownstream end. This penetration increases the effectiveness of thefluid transport to better encase the core stream in the sheath stream.Sheathing will occur, however, even if the fluid transporting structuredoes not penetrate the channel wall. FIGS. 19a and 19b show a twoembodiments of the present sheath flow device having a first fluidtransporting structure 22 located across a top surface 60 of a channeland a second fluid transporting structure 22 located across a bottomsurface 62 of a channel.

Example

The number of grooves can be used to control the position of the corewithin the channel. FIGS. 3(a) through 3(f) show the cross-sectionsresulting from a sheath flow device having 1 pair of grooves through 6pairs of grooves, respectively. One pair of grooves is sufficient tocompletely surround the core stream 28 with sheath stream 26. FIG. 3(a)illustrates the top surface 60 of the channel and the bottom surface 62of the channel. Subsequent pairs carried more sheath fluid to the right,causing the core to be shifted leftward. Having four pairs of groovesappears to be sufficient to place the core roughly in the center of thechannel. Depending on the relative flow rates of the two fluids, thecore can be made as small as 1% of the total channel cross-section. Itis also possible to make the core quite large without losing thesheathing effect.

The fluid transporting structures may also be used in a crossconfiguration when sheath solution is provided from both sides by athird inlet. FIG. 4 shows a channel, 14, having a first sheath streaminlet 10 and a second sheath stream inlet 24. The core stream inlet 12is located between the first and second sheath stream inlets. A firstgroove 22 located in the top of the channel moves sheath stream from theleft of the channel across the top. An opposing groove 22 located at thebottom the channel in a cross configuration with the first groove movessolution from the right of the channel across the bottom. This designhas the advantage that the centroid of the core remains stationary, evenwhen the relative flow rate of the core solution is varied.Additionally, the first and second sheath stream inlets allow differingsheathing materials to be introduced into the channel.

Further, the fluid transporting structures located on the top and bottomof the channel may be configured in a shape that crosses the channelhaving a central area that is distal to its ends, as show in FIG. 5. Thefluid transporting structure 22 of FIG. 4 is shown as a “v” shape,however, any shape having a central area that is located distally in thechannel to its ends would work, such as a semi-circle. FIG. 4 shows achannel, 14, having a first sheath stream inlet 10 and a second sheathstream inlet 24. The core stream inlet 12 is located between the firstand second sheath stream inlets. Fluid transporting structures 22located in the top of the channel moves sheath stream across the corestream to sheathe the core stream.

Example

A microfluidic chip was made using a Techno-isel CNC milling router(Techno Inc., New Hyde Park, N.Y.) in poly(methylmethacrylate) (PMMA)(Plexiglas G, Atofina Chemical Inc., Philadelphia, Pa.) via a methoddescribed by Howell, et al, Lab on a Chip 2005, 5, 524-530, Howell, etal, Lab on a Chip 2004, 4, 663-669, and Mott, et al, Lab on a Chip 2006,6, 540-549, all incorporated in full herein by reference. The mainchannel was 3.18 mm wide by 1.02 mm deep. The grooves were 0.794 mm wideby 0.51 mm deep, and placed in pairs on both the top and bottom of thechannel. A 70% fructose solution was used as core and the sheathsolutions to ensure that the flow within the channel stayed in theStokes regime. The sheath stream was labeled with fluorescent dye(Rhodamine WT, Bright Dyes, Miamisburg, Ohio). Channel cross-sectionsdownstream of the grooves were obtained via a method describedpreviously by Howell, P. B. et al, Lab on a Chip 2005, 5, 524-530 andMott, et al, Lab on a Chip 2006, 6, 540-549, both incorporated in fullherein by reference.

The relative flow rate of the two streams can be widely varied withoutcompromising the integrity of the sheath. FIG. 6 demonstrates acore-to-sheath ratio of 4:1. While the volumetric flow rate of thesheath stream 26 constitutes just 20% of the channel, it stillcompletely surrounds the core stream 28. FIG. 7 demonstrates that acore-to-sheath ratio of 1:4. While the core stream 28 has been reducedto 20% of the net flow compared to the sheath stream 26, it is stillclearly defined. For the specific device and method used in the example,a stable, fully enveloped sheath flow for Reynolds numbers of up toapproximately 200 was generated before the limits of the pump werereached.

FIG. 8 shows a typical cross-section of the channel before sheathing.Sheath stream 26 and core stream 28 are side by side in the channel.FIG. 9 shows the sheath stream 26 surrounding the core stream afterpassing the fluid transporting structures, not shown. Fluorescent dyecan be added to either the sheath stream or the core stream to providecontrast. Unlike other sheath flow systems, this device has also beenshown to be reversible. It is possible to unsheathe a sheathed flow torecapture both the core and the sheath with high efficiency. Unsheathingis achieved by providing a second fluid transporting structure locatedproximally in the channel from the first fluid transporting structure.The second fluid transporting structure is arranged with a reversal ofdirection as compared to the first fluid transporting structure. Thesecond fluid transporting structure does not have to be arranged to bethe exact reverse of the first fluid transporting structure, however,the orientation is in the opposite direction from the first. The abilityto unsheathe a sheathed flow can be useful in systems where the sheathsolution is in limited supply and the capability of recycling the flowis advantageous, such as continuous monitoring on a space station orother enclosed environment. It would also be useful where the solute orparticles in the core solution were very precious and recapture isimportant. FIG. 10 shows the sheath stream 26 and the core stream 28after unsheathing.

The diameters of the sheath and core can vary widely depending on theintended use of the device. FIGS. 8-10 show cross sections of a sheathflow system where the flow rate of the sheath stream is approximatelythe same as that of the core fluid and the sheath and the core havesimilar cross sectional areas. FIGS. 11 a-c show systems in which therelative flow rates of the core stream 28 and sheath stream 26 areadjusted so that the core diameter is very small compared to the sheath(<16 micron core compared to 3 millimeters sheath).

Using specific variations in the pattern of grooves, the exact locationof the core stream can be also be moved across the channel. The capacityeither to separate the walls of the channel from the core fluid using aminimum of sheath fluid or to focus the core fluid in a well definedregion within the channel are significant advantages of the sheath flowdevice and method.

Furthermore, the relative flow rates of the core and sheath can bechanged at will and the diameter of the core can be varied in real timeif the application warrants, with no need to alter the device itself. Asshown in the data in Table 1, the sheathing process remains unperturbed,even at sheath/core ratios over 40,000. FIG. 11a shows a core/sheathratio of 2,100. FIG. 11b shows a core/sheath ratio of 21,000. FIG. 11cshows a core/sheath ratio of 42,000. Higher resolution microscopes wouldenable viewing of fluorescence from the core for even smaller corediameters.

TABLE 1 Reynolds Sheath Sheath Core Core Diameter Ratio of Number FlowRate Diameter Flow Rate Calculated Measured Core/Sheath 0.0008 21 mL/min3 mm 10 μL/min  45 microns 75 microns 2,100 0.0008 21 mL/min 3 mm 1μL/min 4.5 microns  25 microns 21,000 0.0016 42 mL/min 3 mm 1 μL/min  3microns 16 microns 42,000

The actual size of the core can be changed relative to the size of thechannel by simply altering the relative flow rates of the core andsheath streams. Furthermore, this change can be effected in real time.Unlike nozzle system traditionally used for flow cytometry or extrusion,there is no need to go to smaller and smaller nozzles which may resultin clogging problems, higher back pressures, and reduced output. Inprevious designs, the core solution must pass through a nozzle or otherconstriction to enter the flow. This presents a potential cloggingpoint, for the solution containing the cells or other particles to beanalyzed. Under the present design, channels can be of uniform size toavoid constrictions and potential clogging points.

Using the device and method described herein, microdialysis could beaccomplished without a membrane. The core stream is recaptured after itis exposed by sheathing to the sheath stream. This exposure provides forthe removal of low molecular weight molecules by diffusion across theinterface of the core stream and the sheath stream. The ability toconduct microdialysis without a membrane prolongs the life of thesystem. Current microdialysis systems operate for limited lifetimes dueto the potential for membrane clogging. Additionally, separations basedon differential solubility as well as differential size can be providedby the device and method described herein. For example, a whole bloodsample could be sheathed into the center of the channel, and allowed toflow for sufficient distance for small molecules to diffuse outward fromthe core into the sheath. Cells and larger molecules such as proteinswill not diffuse as quickly and will tend to stay in the core. The corewould then be unsheathed and recovered, with the smaller moleculesremoved.

The device and method are useful as a means of protecting conduits,including but not limited to, pipes, tubes, ducts, tubing, capillaries,and microfluidic channels, from fouling or corrosion. A thin sheathstream of protective material is formed around the core stream. Thesheath stream need not be the same viscosity as the core stream,therefore a relatively slow moving and thin protective sheath coatingcan be formed to protect the insides of conduits exposed to corrosivecore stream solutions.

The device and methods described herein can also be used to reduce thepower requirement for transporting viscous fluids in conduits, includingbut not limited to, ducts, pipes, tubes, tubing, capillaries, andmicrofluidic channels. Sheathing a viscous fluid in a second fluid oflower viscosity reduces the sheer stress at the conduit wall whichlowers the pressure drop required to generate a given flow rate. Thesheath flow component has been used to generate such a flow, in which acore and a sheath stream of differing viscosity initially enter thedevice side-by-side and the lower viscosity sheath stream sheaths thehigher viscosity core stream.

The relatively low flow resistance of the device means that it can beused to sheath quite high-viscosity systems. This is useful in food andpolymer extrusion applications. The device and method is further usefulin the synthesis of specialty polymeric filaments and tubes. Unlikestandard extrusion technologies, filaments with continuously varyingdiameter can be created. Filaments made in this way can be expected tohave increased elasticity over extruded filaments because of the nativeentropy of the polymer chains. The exact design may also be altered tochange the cross-sectional shape of the resulting polymer strand. Sincethe extrusion device is small, inexpensive, and essentially operates asa passive component, many devices can be fabricated to perform inparallel, such as an array.

The device and methods described are also useful as liquid waveguides.Liquid waveguides have been described for monitoring chemical processesin which light is guided in fluid in a capillary or in the walls of acapillary in order to measure some component of the fluid. The deviceand method can be used for guiding the light in either the core streamor sheath stream for similar measurements, but with the capability formore exact focusing, much greater control of the relative dimensions ofthe light guiding fluid and the other fluid, and the avoidance of walleffects such as scattering of the light from the core by the capillarywall. The capability of guiding light in fluids is particularly usefulin microfluidic systems.

FIG. 12 shows the waveguide application. A chip 31 was fabricated with achannel 14 beginning in the center and spiraling outward to the outlet20 on the outside edge of the chip. A sheath stream inlet 10 and a corestream inlet 12, located near the center of the chip, are in fluidconnection with the channel 14. The fluid transporting structures 22sheathe the core stream within the sheath stream. The sheathed solutionthen travels outward in a spiral of 360 degrees before reaching theoutlet 20. A light source 32 is introduced through a window (not shown)located at the outlet 20.

Core and sheath streams are introduced into the structure at the inlets.The core and sheath streams have approximately equivalent densities. Thecore stream is 70% fructose. The sheath stream is a saturated saltsolution with enough fructose added to match the density of the core.There is a small amount of fluorescent dye in the sheath stream. Thesheath was formed in the center of the chip 31 and then traveled outwardalong an increasing spiral.

FIG. 13 shows the resulting waveguided light 33 when light wasintroduced to the channel from an outlet 20. The light is waveguided 33through a full 360 degrees around the spiral. The light source 32illuminates the higher refractive index stream, which in this case isthe core; however, it could be either the sheath stream or the corestream.

The condition for waveguiding is merely that the core stream and thesheath stream have different refractive indices. The ability tohydrodynamically focus a core down to submicron diameters allows for theproduction of a nearfield optical microscope probe entirely out ofliquid. FIG. 14 shows an example of a nearfield optical microscopeutilizing the present invention. Once the core stream 34 is ensheathedin the sheath stream 36, a tapered nozzle 38 is used to create the taperin the core. The high refractive index core stream 34 is directedthrough the nozzle 38. Light introduced into the core will be waveguideddown to the surface 40. Reflected, scattered, or emitted light will thenbe collected by the waveguide and carried upward for detection. Anotherpossible design may eliminate the need for a nozzle by introducingdielectrophoretic forces to push the core stream out into a fine tip.This design would also be able to use dielectric forces to steer thestream and raster it over the surface. Based on refractive indexmeasurements of the chosen chemistry, the optimal geometry of the tapercan be established. Because a solid tip does not have to be brought intoclose proximity with the surface, this design is well suited for theanalysis of fragile biological samples. It is also well suited toperform liquid-phase photochemistry for nanomachining processes. Thechip is able to raster over a surface using a translation stage.

FIG. 15 shows an exemplary flow cytometer using sheath flow. The inlets10, 12 are connected to the channel 14. The fluid transportingstructures 22 wrap the sheath stream around the core stream, focusingthe core stream in the interrogation region 46. Interrogation, forexample, illumination, comes from a single mode fiber 42. Light wascollected by a multimode fiber 44.

FIGS. 16(a)-16(e) show a series of traces of the light scatter resultingfrom the series in order of increasing concentration. As shown in thesignal tracings, representing the light scatter signal from five-foldserial dilutions of yeast cells, the light scatter signals wereproportional to the concentration of cells in the flow stream. Thesample core was illuminated with the light from a helium-neon laserintroduced via a single mode optical fiber. Scattered light wascollected at 90 degrees using a multimode fiber and detected with aphotomultiplier tube. FIG. 16a was a highly diluted sample, showing nocells during the 4-second sampling time. Each successive solution wasroughly 5 times as concentrated as the previous solution. Each of thespikes seen in a plot represents the passage of a cell through theinterrogation region. The number of spikes increases approximately5-fold with the 5-fold increase in concentration.

The device and method of the present invention are also useful for thefabrication of materials. For example, the core stream can contain apolymerizable, condensable, cross-linkable or crystalizable material,which is extruded to the desired diameter using the sheath streaminstead of a solid nozzle or channel. Since the sheath flow device issmall, inexpensive, and essentially operates as a passive component,many devices can be fabricated to perform in parallel, such as an array.

Materials from which fibers or other structures can be fabricatedinclude but are not limited to a wide variety of polymers includingpolystyrene, butyl rubber, polypropylene, polyacrylamide, polysiloxane,and polymethylmethyacrylate. Biological molecules can be ordered toself-assemble into higher order structures; such molecules could includea wide variety of lipids, proteins, carbohydrates and oligonucleotides.Materials that form harder structures could be used including precursorsof glassy materials such as sol gels, as discussed in Sousek et al.,Polymers for Advanced Technologies, 2005, 16:257-261, incorporatedherein in full by reference, or initiators for subsequent deposition ofmetals, calcium, and/or semiconductors. The fluids used can be aqueousor organic. Preferably, the core and sheath fluid are the same phase.

By varying the diameter of the core, tapered materials can befabricated. Nonuniform or tapered geometries for waveguides can begenerated. Controlling the relative rates of sheath and core flow duringpolymerization of filaments provides high precision, tapered structureswith sub-micrometer diameter fluctuations, resulting in uniquewaveguiding properties.

The device and method is further useful in the synthesis of specialtypolymeric filaments and tubes. Unlike standard extrusion technologies,filaments with continuously varying diameter can be created. Filamentsmade in this way can be expected to have increased elasticity overextruded filaments because of the native entropy of the polymer chains.The exact design may also be altered to change the cross-sectional shapeof the resulting polymer strand.

By configuring the grooves or ridges used to transport the sheathstream, non-round shapes can also be obtained. In addition to varyingthe rate of flow to change the diameter of the core, the core fluid canbe pulsed instead of flowed continually to stop and start the corestream to form “particles” or “packets” of core fluid. Once the desiredsize and shape are obtained, the material in the core is polymerized,condensed, cross-linked, or crystallized chemically, optically or byother means known in the art. Due to the geometry of the system, thistype of synthesis can be conducted in continuous manner rather than inbatches. Moreover, the geometry of the system is particularly amenableto the production of high-aspect-ratio structures and filaments that areespecially difficult to produce in quantity.

Shapes that can be fabricated in this method include, but are notlimited to, ovals, ribbons, rods, wires, tubes and filaments. Using thegrooves or ridges on the top and bottom of the channel can bespecifically designed to produce the desired shape. The grooves orridges do not have to be straight but can have a variety ofconfigurations as long as they channel the fluid around the core. Theycan be curved, in the shape of chevrons, angled like “check marks,” orin a variety of other shapes in order to influence the shape of theresultant core fluid. The addition of more inputs and grooves furtherdownstream can be used to expand the repertoire of shapes that can befabricated.

More complex shapes that can be designed and fabricated using grooves orridges include hollow cylinders, filled “sausages,” coated particles,rods with alternating composition, also known as “nano bar codes”.Structures with longitudinal or lateral density or chemical gradientscan be fabricated by introducing gradients into one of the flow streams(longitudinal) or by allowing a reactant to diffuse in or out of thecore while it is in contact with the sheath stream (lateral).

FIG. 17 shows a sheath flow device capable of creating a hollow tubeswithin a hollow tube. A sheath input 10 and a core input 14 areconnected to a channel 14 having a series of fluid transportingstructures 22. A series of successive sheath inputs 10 are provideddownstream towards the outlet 20. Each successive sheath input 10creates a new sheath around all the previously sheathed materials. FIG.18 shows a hollow tube within a hollow tube that was made by using thedevice of FIG. 17 by alternately introducing input streams andensheathing the structures. The deepest interior core stream 50 issurrounded by successive core/sheath streams 52, 54, 56, followed bysheath 58. Streams 52 and 56 were labeled with a fluorescent dye forcontrast.

Typically, the sheath stream is sufficient to move the polymerizedmaterial to the output of the channel. For some materials, however, asthe extruded material polymerizes and its viscosity increases from itsunpolymerized value to infinity, the dynamics of the flow profile withinthe channel may change to the point that feed matching is required tocontrol the fluid velocity and effectively remove the polymerizedmaterial. There are several options available to do feed-matching. In anelastomeric chip, the fluid velocity is controlled by compressing thechannel to cause the fluid to accelerate. Additionally, rollers may beplaced at the exit of the chip so that they impinge on the rod andcontrol the linear exit velocity of the polymerized rod.

Generally, the core contains a polymerizable material and is extruded tothe desired diameter using the sheath stream instead of a solid nozzleor channel. Once the desired shape is obtained, the core material ispolymerized chemically or optically. Due to the geometry of the system,production can be in continuous instead of in batch mode. Moreover, thegeometry of the system is particularly amenable to the production ofhigh aspect ratio structures and filaments which are especiallydifficult to produce in quantity. Since the fabrication device is small,inexpensive, and essentially operates as a passive component, manydevices can be fabricated to perform in parallel, such as an array.

Multiple Fibers Fabricated in a Single Channel.

Structures, including but not limited to grooves, ridges, and pillars,can be made in the channel floor and roof such that multiple fibers canbe produced in a single channel configured to produced multiple sheathedflows. The core streams in the multiple sheathed flows are polymerizedto form multiple fibers therefrom. This can be accomplished by shapingmultiple inputs of polymerizable material as in FIG. 20 or by splittinga single prepolymer stream into parallel streams as in FIGS. 21A and21B, which show a simulation and actual splitting, respectively, of afiber caused by controlling both grooves in the top and bottom of thechannel and the relative flow rates of the sheath and polymerizablematerial. By changing the flow-rate ratios, fibers could be fabricatedwith split regions that recombine into a single fiber. This could bedone repeatedly to obtain individual fibers with repeated split regions.FIG. 23 exemplifies a top-down view of a design that could be used tocreate multiple parallel fibers.

In embodiments, a split fiber has two lobes that remain attached, or thesplit may be complete so that two or more distinct fibers arise, such asfrom a corresponding number of distinct core streams.

Production of Fibers with Dimensions Altered Along Length of Fiber.

Relative flow rates of the polymerizable material(s) and sheathsolution(s) can be altered during the polymerization/casting process tocreate fibers with variable dimensions along the length of the fiber. InFIG. 24, variations in the pump speed changed the ratio of flow rates ofpolymerizable material and sheath to produce a fiber with across-section that alternately thickened and thinned. The ability totaper a fiber could be especially useful for optical applications.

Production of Particles, Rods and Packets.

Modifications in the methodology for producing shaped fibers can be usedto produce particles, rods, or packets in the same type of microfluidicdevices. A packet refers to an enclosed hollow shape. The method of usecan generate control over the length of the structures by a variety ofmechanisms while the cross-sectional shape is determined as alreadydescribed for fibers. Mechanisms for breaking the continuity of the corestream can include “chopping” the light used for UV polymerization tomake rods or particles with defined length; subjecting the core topiezoelectric, acoustic, or other alternating forces to move the coreback and forth in the sheath stream; alternating polymerizable andnonpolymerizable chemical solutions in the core stream; and/or usingvariations in flow-rate ratios to pinch off the core into discreteparcels.

Similarly, magnetic or electric forces could be applied in a fixed ormodulating fashion across a channel in order to modify the alignment ofmaterial, including the polymerizable material and/or content suspendedtherein. Such forces can also be used to create a gradient, discussedbelow.

Electric or magnetic fields can also be applied longitudinally along thechannel (parallel to the direction of flow) to encourage alignment ofpolymer chains. For example, with cationic polymerization, alongitudinal electric field could draw the positively-charged reactioncenters in a direction along the forming fiber, causing the polymerbackbone to trail behind.

Production of Fibers with Encapsulated Cells, Enzymes or OtherBiological Elements.

Fibers or packets could be made with biocompatible polymers, includingbut not limited to collagen, agarose, polyelectrolytes, chitosan,gelatin, polyethylene glycol, or peptides, and derivatives of thesemolecules, and used as matrices, scaffolds or hollow supports for tissueengineering or extended cell culture for cells from mammals or otheranimals. Hollow fibers or packets could include cells (such as mammaliancells), the cells could grow on inner or outer surfaces of the fiberswith nutrients and/or antibiotics delivered through the fibers, or cellscould be embedded in the fibers. A significant advantage of thistechnique is that the cells can be introduced during the manufacture ofthe fibers, in the lumen, in a biocompatible prepolymer layer, or in thesurrounding sheath fluid. Alternately, a conventional approach of makingthe scaffold structure and introducing the cells thereafter is alsopossible. The applications for such materials include, but are notlimited to, medical and scientific research, wound healing, tissueengineering, pharmaceutical screening, and bioprocessing.

Fibers could be used to encapsulate cells (such as bacterial cells) orspores selected or engineered for biomanufacturing, biosensing, orbioremediation. For use in the field, whether as sensors or fordecontamination, cells must simultaneously be protected from theenvironment and exposed to it. The level of tolerance to non-optimalconditions is much higher for bacteria genetically modified for sensingor selected for degradation capability than for animal cells, though thelatter are also under development. Bacteria have been encapsulated,immobilized, or used free in solution. The first approach usuallystabilizes the bacteria but can limit transport of the target compoundto the bacteria, the second often damages the bacteria, and the thirdrequires large quantities relative to the fluid being tested. Testingtimes range from hours to days, depending on the resistance of thebacteria to the toxicity of the sample matrix.

The inclusion of target-reactive bacteria in hollow fibers, along withnutrients and stabilizers (e.g. trehalose), can be used for continuousmonitoring of effluents from air samplers, drinking water, or othersources. Cell lines reported in Anal. Chem., 82: 6093-6103 (2010) areexemplary candidates for such use. These cells form spores that arehighly stable for long periods (24 months at room temperature or 12months under extreme temperature and humidity/drought environments), yetcan be germinated and produce a measureable response to target analytesin ˜2 hours. The two genetically modified lines generate a luminescentsignal in the presence of zinc (Bacillus megaterium) or arsenic(Bacillus subtilis). The zinc sensing system employs the enhanced greenfluorescent protein (EGFP) as a reporter, which is detected by excitingwith UV light, while the arsenic sensing system utilizesβ-galactosidase, which can be detected by a chemiluminescent substrate.Substrates are present within the spores and do not need to be addedexogenously. Spores are ideal biosensing elements in that they arerugged, inexpensive to produce and easy to make and germinate. Indeed,sensing spores can be cycled from dormant to active over a period of atleast two years without any significant loss in their analyticalperformance. Moreover, storage of spores under a variety of stressfuland stringent conditions does not affect their sensing ability whenbrought back to active cells.

Since the conditions for shaping the fibers with hydrodynamic focusingare typically performed with relatively low sheer stress and thepolymerization conditions are generally mild (casting or brief exposureto UV light), the fibers can be formed with cells present in one of moreof the fluids. For incorporation of cells into the polymer layer, thecell-containing fluid, including stabilizers if necessary, and hydrogel,protein, and/or other prepolymer (preferably biocompatible) is focusedwith a sheath solution that is also preferably biocompatible and ofmatching viscosity. Microchannels with grooved structures in the top andbottom of the channel can be designed to use a phase-matched sheathfluid and were demonstrated to focus a polymerizable core stream into apredetermined shape without mixing. The cross-sectional dimensions canbe determined by the relative flow rates of the sheath and core. Usingmultiple fluid additions, successive layers of fluids can be wrappedaround the core. The cell-containing core fluid, including stabilizersif necessary, and a hydrogel or other biocompatible polymer is definedso that it matches the viscosity of the fiber prepolymer. Thecomposition of the fiber prepolymer is designed for rapid polymerizationwithout cell damage; in addition to the acrylate recipes used to date,we can use polymers based on click chemistry (e.g. Applications of clickchemistry themed issue of Chemical Society Review, edited by M. G. Finnaand V. Fokin 2010, especially C E Hoyle, A B Lowe, and C N Bowman Chem.Soc. Rev, 2010, 39:1355-1387). Composition and thickness of the hollowfiber layers can be adjusted as necessary to provide strength andstability with maximum diffusive transport.

In experiments designed to demonstrate fabrication with cells, bacteriawere included in a hydrogel prepolymer and bio-hybrid fibers werefabricated using hydrodynamic focusing as described in Daniele M A,North S H, Naciri J, Howell P B, Foulger S H, Ligler F S, et al. “Rapidand continuous hydrodynamically controlled fabrication of biohybridmicrofibers,” Advanced Functional Materials. 2012, 23:6 698-704. Thecells were viable after the polymerization of the fiber and capable ofcontinued proliferation in the fiber.

In another experiment, hollow fibers were made using hydrodynamicfocusing. FIG. 30 shows a device that focused first a nonpolymerizablecore fluid with a polymerizable sheath fluid (introduced from a singlereservoir into inlets on both sides of the core inlet). Fromleft-to-right, a core fluid is introduced and hydrodynamically focusedvertically in the center of the channel by the first sheath fluid. Thefluids traverse the chevrons that serve to: (1) compress the core fluidvertically and (2) symmetrically wrap the sheath fluid about thecompressed core. This concentric two-component flow stream thenencounters a second sheath fluid, whereby it is focused againhydrodynamically in the vertical direction while being encircled in thesecond sheath fluid after traversing the second set of chevrons. Themicrofluidic channel that follows serves as a reactor wherephotoinitiated polymerization is carried out, locking in the shape ofthe polymer. The core fluid and first sheath fluid can be independentlypolymerizable materials yielding either a two-layer/component solidfiber or a two-layer/component liquid core solid shell fiber which isequivalent to a hollow fiber. Also, it is important to note that thefinal sheath used to protect the concentric flows from contact with themicrochannel walls can be doped with reactive species complementary tothe chemistry used in the outermost shell or in the lumen to formadditional thin polymer layers that correspond to the diffusion boundaryof the photoinitiator containing outer shell. This approach might serveto add a skin-like layer to enhance integrity or mimic the membranesaround (for example) muscle fibers. Such an embodiment is also describedbelow under “Production of fibers via interfacial reactions.”

In a second shaping region, the first set of core and sheath streamsessentially traveled down the channel as the core fluid and wereensheathed with a nonpolymerizable fluid of equivalent viscosity. Inthis experiment, hollow fibers or tubes were produced, as shown in FIGS.31A and 31B, with air bubbles passing through the lumen for easiervisualization of the inside of the hollow fiber. Hollow microtubes wereproduced with diameters on the scale of arterioles and venules, i.e.“capillary-sized” fibers.

A cell-fiber system incorporating bacteria is preferably designed toaccomplish the following:

1. The cells are stabilized in a ready-to-use format during shipping andstorage at room temperature.

2. Nutrients needed for cell reactivation are encapsulated into thecell-fiber mat for operator convenience using several strategies.

3. The cell-fiber mats provide high surface area-to-volume for sampleinterrogation.

4. The cell-fiber mats provide a convenient footprint for automated,continuous monitoring.

5. Optics for luminescence detection are very simple (filter andphotodiode) and can be battery operated.

6. Cells with new specificities for detection, catalysis or degradationcan be genetically engineered and incorporated into the fibers, eithersingly or in mixtures.

7. The fiber mats prevent release of genetically modified organisms intothe environment. Used materials can be easily destroyed for safedisposal.

The technology developed for sensing can be extended to decontaminationwith the availability of appropriate cells. The fibers can be aligned orwoven to make filters for decontamination or textiles for protection ofwarfighters or hazmat workers.

Fibers or Particles with Encapsulated Enzymes or Other BiologicalElements.

The embedding of active biomolecules (such as enzymes or other proteins)in the fiber is simpler than encapsulating active cells. Methods forencapsulating active enzymes in hydrogels, sol gels, polymer beads,polyanionic films, and other materials are well documented.Nevertheless, there is still a need for maintaining biomolecularactivity in filters, woven fabrics, beads, and other solid phases usedfor biomanufacturing, separations, remediation, protection, and sensing.The active biomolecules can be encapsulated randomly throughout theshaped fibers, along with any required stabilizers or cofactors, orthese molecules can be included in a core layer surrounded by a layerpolymerized to have the optimum porosity for the desired function.Molecules that promote capture and transport of the target from theoutside to the inside of the fiber can be included throughout or just inthe outer layer(s).

For example, the fibers could include a polymer matrix of appropriateporosity and containing carboxylic moieties, Cu²⁺ chelated to the vinylgroups for binding phosphonates, and an enzyme for catalysis. It hasbeen demonstrated that hydrodynamic focusing in microfluidic channelscan be used to fabricate porous acrylate fibers with pre-designedcross-sectional shapes (see Thangawng et al., Lab Chip 9 (2009)3126-3130). Round fibers or flat ribbons have been made with dimensionsfrom ˜300 nm to ˜300 μm in lengths up to meters. The fibers have beenspooled so that they are aligned in parallel or collected them inrandomly organized mats. The fibers have been characterized in terms ofshape, dimensions, molecular organization, and tensile strength.Depending on the size and method of polymerization (casting or UV), thefibers can make them more or less porous. A key metric will be theamount of liquid or air that can be wicked into a gram of fiber, whichwill depend on the fiber geometry, weave pattern (pores, capillaryaction), and fiber chemistry (surface wetting, swelling, internalporosity). Shape and organization are important since as the distancethat the toxic agent must diffuse from the surface of the fiber to theactive components is decreased, the faster the target molecule can bebound and/or degraded.

The same considerations apply to the encapsulation of molecularrecognition elements, with or without enzymatic activity. Once coulddesign fibers or particles that include sensing molecules along withmolecular elements for signal generation, including but not limited tofluorescence, chemiluminescence or electrochemical signals. The responseto molecular recognition could include controlled release of a drug ortherapeutic, such as the release of insulin in response to detection ofhigh glucose levels in vivo.

Shaping of Fibers for Assembly into Larger Scale Materials with New orImproved Properties.

Round or non-round fibers may be formed into larger scale materials.Exemplary larger scale materials include, for example, textiles,composite films, environmentally sensitive smart materials, highstrength materials, cables, yarns, etc. Fibers produced by standardmethods such as extrusion or electrospinning are round due to theminimization of interfacial tension at the boundary between theprepolymer core and surrounding air or other phase, however theproperties of non-round fibers may be exploited in larger scalematerials. For instance, post-polymerization modification, such astwisting, of non-round fibers can create periodic structures in thefiber. Larger scale materials can also be prepared by techniques knownto those of skill in the art, for example, spinning, weaving, and/ornonwoven production methods (staple nonwovens, spunlaid nonwovens).

The interaction of shaped filaments or fibers in such materials couldprovide new or improved strength, flexibility, potential for actuation,or other new properties. For example, alignment of many small filamentswithin a rope or textile leads to improved strength compared toindividual large fibers. Various types of fiber shapes and exemplarylarger scale materials are illustrated in FIG. 25.

Phase-separating materials, such as certain polymer blends or blockcopolymers can be used to cause the self-assembly of structures withinor on the surface of the fibers. These structures may be aligned with oracross the fiber and can play a role in the formation of larger scalematerials.

Production of Fibers Via Interfacial Reactions.

The polymerization, precipitation, or other hardening reactions can beinitiated by the combination of compounds that takes place at thecore/sheath boundary, or other interface between streams such as betweenan interior core stream and exterior core stream. It is also possible tohave multiple streams that come in contact to produce a reaction. Suchtechnology could be broadly classified as an interfacial reaction.

The kinds of interfacial reactions that can produce polymers duringhydrodynamic focusing can generally be classed into two categories. Inthe first, the reaction that takes place at the interface immediatelyproduces a solid product, which ultimately seals the interface and capsthe reaction so that the reaction is limited to the interface. Theresult would be an extruded material at least initially in the shape ofthe interface. One example of such a reaction is in the production ofnylon, such as at a hexane/water boundary. The second category ofreaction is one which can continue to propagate into the bulk of thematerial once the two flows are brought into contact, for exampleprecipitation of PMMA described below, wherein the PMMA solvent is stillmobile in the solidified PMMA, and continues to leave the fiber evenafter the perimeter has already hardened. Another example would be theintroduction of an initiator to a living polymerization. Once initiatedat the interface, the reaction center can then continue to migrate intothe bulk of the liquid monomer.

In an example of an interfacial reaction propagating into the bulk ofthe material, a solution of polymethylmethacrylate (PMMA) in acetone(other suitable solvents can be used) was sheathed in an aqueoussolution. As solvent diffused out of the core into the sheath, the PMMAprecipitated to form a fiber. In addition to simple precipitation, otherreactions can be used, including acid/base chemistry, introduction ofchemical initiators, and step-growth polymerization.

Although the shaping grooves function most reliably in one-phasesystems, fibers have been produced in two-phase systems as well. It isalso possible to temporarily remove the interface of a 2-phase system byplacing a thin layer of an intermediate solvent between the twomaterials. As an example, a thin layer of isopropanol (IPA) can beplaced between water and hexane. Being miscible with both water andhexane, the IPA will replace the sharp water/hexane boundary withdiffuse water/IPA and IPA/hexane interfaces. If made of appropriatethickness, the IPA will maintain the one-phase condition through theshaping of the fluids, then diffuse away sufficiently the water/hexaneinterface to be reestablished.

Production of Fibers with Lateral Variation in Composition.

More than one polymerizable material can be incorporated into the samecore, resulting in fibers with multiple compositions in a predefinedconformation. FIG. 18 shows a cross-section through a channel whereseveral concentric flow streams have been created. Inclusion ofpolymerizable materials in the flow stream would produce a fiber withseveral concentric layers. A concentric configuration could beparticularly useful in situations where a fiber is wanted with differingbulk and surface properties. It is expected that a wide variety ofconfigurations of two or more polymerizable materials can be constructedusing the shaping structures. FIG. 26 shows just a few possibleconformations that can be created. The lateral composition does not haveto be discrete. Elements placed upstream of the sheathing can bedesigned to create continuous gradients as well, such as those found ingradient index fibers. Alternatively, merger of multiple streams priorto polymerization can be used to create lateral gradients.

Gradients can exist in one or more components of the core and/or sheath.For example, gradients can exist in the concentration of crosslinker,ions, and/or polymerizable material. More than one gradient can existsimultaneously.

Not all of the structures in the shaped streams need to be somethingthat ultimately hardens. For example a hollow fiber could be made thatis filled with a liquid. Because structure can also be changedlongitudinally, the lumen can be pinched off periodically, so that a cutin the fiber does not cause its entire length to drain. Applications forthis kind of structure would include drug release, contaminantsequestration, phase-change thermal fabrics, etc. The structures couldalso be deliberately drained after fiber production, thereby creatingvoids. Large voids could be used as tubing, while multiple smaller voidshave a possible application in photonic materials. It should also benoted that unlike classical extrusion technologies, the voids can bemade to split or recombine by the same mechanism used to make fibersthat split and recombine (as noted above with regard to multiple fibersfrom a single channel).

Production of Fibers with Residual Stresses (Curly Fibers).

Another layer of structure can be added to the fiber by introduction ofresidual stresses. Many polymers contract during curing. By using theability to make fibers of differing lateral compositions, one coulddeliberately engineer the contraction to cause the fibers to curl.Differential curing could also be induced by chemical, light or othergradients. FIG. 27 shows a fiber where stresses were introduced, mostlikely due to a light shining on one side. Prestressed fibers are notlimited to round shapes. A fiber could be produced with a central coreand one or more long “wings” extending from the core, seen in FIG. 28.If the core is designed to contract during curing, the wings willdevelop a scalloped or frilled pattern. Such a fiber would have a highersurface area per unit length, making it well suited for filtering orcatalytic applications.

Deliberate Buckling, Breaking, or Other Effects of Applying Forces onNascent Fibers.

As the material comprising the fiber hardens, it can be subjected toforces that affect its ultimate molecular or gross structure. Oneexample can be seen in FIG. 29. When a viscous fluid stream is forced todecelerate, it can buckle upon itself as seen here. This viscousbuckling can take other forms, including the spiral motion of a streamof syrup as it lands on a surface. A similar behavior could be used tocreate helical fibers. The forming fibers could also be exposed tostretching or bending. If the material is ductile enough, this couldsimply help to align polymer fibers or have other desirable effects onthe composition of the fiber. If a more brittle material is used (e.g.sol gels), the result could be the break-up of the fiber into rods ofregular size, shape, and aspect ratio.

Printing with Sheathed Flow.

A partially polymerized stream can be directed onto a surface while thesurface is moved relative to the stream, or vice versa. Ifpolymerization is incomplete at the time the fiber is laid down, thereis a tendency for the fiber to adhere and conform to the substrate onwhich is it being laid. Preferably under computer control, patterns canbe laid down on a surface, and multiple levels of a fibers could be laiddown to print three-dimensional objects, with the resolution of theobject set by the diameter of the fiber.

Tissue Engineering with Animal Cells

Hydrodynamic focusing using core/sheath flow as described above may beused to create polymer fibers containing cells from mammals or otheranimals. The fibers can be comprised of one or more layersconcentrically arranged around a central lumen (derived from the corestream), which is optionally hollow or filled. In the case of a hollowlumen, the core stream would include material that is not polymerized,whereas including polymerizable material in the core stream could createa lumen filled with polymer to form a “solid” fiber. The cells can beinside the lumen, embedded in one or more layers, and/or adherent to theoutside surface of the fiber. The cells in each layer may be the same ordifferent in origin, with any type of cell found in an animal acandidate for use in this technique, including cells from mammals,birds, fish, reptiles, amphibians, insects, etc. The materials in thepolymer are preferably biocompatible. Other components beyond cells andpre-polymer can be included in the flows and thus be incorporated in theresulting fibers, for example factors to modify cell growth, adhesion,and/or differentiation. It is also possible to incorporate nucleic acids(such as RNA or single or double stranded DNA fragments) in the polymermatrix. These nucleic acid moieties could be used to bind proteins orpeptides or as templates for replication or functionalization.Furthermore, materials can be selected so that 90% or more of thepolymer in the fiber is degraded by the cells and replaced withextracellular matrix, or only certain of the concentric layers may bedesigned for such biodegradation. The cell-containing fibers can mimicnatural structures such as capillaries, blood vessels, tissue ducts, ornerves. Thickness of individual polymer layers is limited primarily inorder to allow for diffusion of oxygen and nutrients to the cells.

Microchannels with grooved structures in the top and bottom of thechannel can be designed to use a phase-matched sheath fluid to focus apolymerizable core stream into a predetermined shape without mixing, asdescribed in members of this patent family and publications of inventorsthereof, including Boyd, D. A., A. R. Shields, J. Naciri, and F. S.Ligler (2012) “Hydrodynamic shaping, polymerization, and subsequentmodification of thiol click fibers,” ACS Appl. Mater. Interfaces,recently accepted manuscript, DOI: 10.1021/am3022834; Shields, A. R., A.L. Thangawng, C. M. Spillmann, J. Naciri, P. B. Howell, and F. S. Ligler(2012) “Hydrodynamically directed multiscale assembly of shaped polymerfibers,” Soft Matter 8, 6656-6660; Thangawng, A. L., P. B. Howell, C. M.Spillman, J. Naciri, and F. S. Ligler (2011) “UV polymerization ofhydrodynamically shaped fibers,” Lab Chip 11, 1157-1160; and Thangawng,A. L., P. B. Howell, J. J. Richards, J. S. Erickson, and F. S. Ligler(2009) “A simple sheath-flow microfluidic device formicro/nanomanufacturing: fabrication of hydrodynamically shaped polymerfibers,” Lab Chip 9, 3126-3130. The cross-sectional dimensions can bedetermined by the relative flow rates of the sheath and core. Usingmultiple fluid additions, successive layers of fluids can be wrappedaround the core, as shown in FIG. 18.

For use of animal cells during the fabrication process, it is desirablethat any photopolymerization of the prepolymer material be accomplishedrapidly (with minimal if any damage to the cells) and that the flows beunder moderate pressure (to reduce shear stresses on the cells). Theengineered material used in the experiments described here was abio/synthetic hydrogel composed of gelatin methacrylamide (GelMA) andpoly(ethylene glycol) (PEG). By modifying the gelatin with methacrylategroups and PEG with thiol or alkyne groups, a biocompatible hydrogel wasformed that can be rapidly crosslinked by mild UV exposure and hadadjustable physiochemical properties. The GelMA was selected for thepresence of the peptide RGD known to exhibit binding sites thatfacilitate cellular ingrowth and extracellular matrix formation. The PEGwas chosen to provide mechanical support during formation of the hollowfibers and matrix remodeling (e.g. cell-induced degradation of thegelatin and replacement with extracellular matrix). The hydrogel provedto be biocompatible for both cell adherence and encapsulation ofendothelial cells. Other biocompatible materials can also be used aslong as the conditions for polymerization or casting do not damage thecells.

Polymers may include proteins, peptides, collagen, agarose,polyelectrolytes, chitosan, gelatin, hyaluronic acid,heteropolysaccharides, polyethylene glycol, hydroxyethylmethacrylate(HEMA), gelatin methacrylamide (GelMA), poly(methyl methacrylate)(PMMA), poly(lactic-co-glycolic acid) (PLGA), polylactic acid (PLLA),derivatives of any of these, and combinations thereof.

By utilizing an additional set of hydrodynamic shaping features in thetop and bottom of the channel as seen in FIG. 32, the biocompatiblehydrogel precursors are directed into concentric layers andphotopolymerized in situ. To make a two-layer hollow tube, the first“core fluid” is not polymerizable. The initial stage of the devicefocuses the nonpolymerizable core, while the second stage focuses theprepolymer designed to become the inner layer of the hollow fiber.Unlike the example depicted in FIGS. 30 and 31, however, the sheathfluid introduced in the second stage is also polymerizable. In the thirdstage, the first three solutions are completely ensheathed in a fourth,nonpolymerizable solution. The concentric polymerizable fluids are thenpolymerized after passing through the final set of fluid transportingstructures, preferably using broad-spectrum ultraviolet light (≤10mW/cm²). The diameter of the lumen and the thickness of the layers inthe hollow fiber is a function of the relative fluid flow rates of eachof the fluids introduced into the microfluidic channel. FIG. 33 showscomputational fluid dynamic models of the lumen and the polymerizablelayers that would be obtained using variations in shaping features andin the flow-rate ratios and cross-sectional areas of flows produced fromtwo sets of chevron shaping features, three sets of chevron shapingfeatures, and two sets of striped shaping features. The flow-rate ratiosare consistent in all the shaping simulations. Exemplary flow-rateratios that provided good results were core:sheath_(n):sheath_(n+1) suchthat each subsequent layer is 4 times the flow rate of the previouslayer. For example, a core flow with 3 sheaths would have flow-rateratios of 1:4:16:64 using this formula, however variations are possible.

Cells can be included in any of the streams during the fabricationprocess, depending on the desired configuration of the model system.Blood vessels include endothelium on the inner layer, smooth musclecells in the primary wall of the vessel, and a surrounding layerincluding fibroblasts and extracellular matrix, as depicted in FIG. 33.During the fabrication of a model blood vessel, endothelial cells areincluded in the first nonpolymerizable core fluid, smooth muscle cellsare included in the prepolymer used as the first sheath fluid, andfibroblasts are included in the second sheath propolymer fluid. Thenonpolymerizable fluids included cell culture media with unmodifiedgelatin or PEG to increase the viscosity and match it to the viscosityof the polymerizable fluids. The bio/synthetic hydrogel employed toproduce the multi-walled microtubes is seeded with cells beforeinjection into the microfluidic shaping device. The luminal fluidcontains Human Umbilical Vein Endothelial Cells (HUVEC) that can attachto the interior surface of the micro-blood vessel. The first prepolymerfluid for making the interior elastic layer incorporates PrimaryCoronary Artery Smooth Muscle Cells (PCSMC). The prepolymer for theexternal layer incorporates Human Dermal Fibroblasts. The shapingfeatures focus the cell-containing flows into the respective zones toproduce the multi-walled microtubes. Polymerization is sufficientlybrief that cell damage does not occur. The generated micro-blood vesselsare collected directly into sterile cell culture medium for individualstudy or for incorporation into tissue-on-chip model systems.

This microfabrication approach is modular, so that additional inlets andshaping features can be appended to the design to produce any number ofconcentric polymer layers. The previously tested GelMA-PEG hydrogel hasbeen used to form the microtubes and has proven biocompatibility andmechanical strength. The flexible photochemistry also allows for theinclusion of instructive biological components, a popular methodology intissue engineering to encourage cellular proliferation. [Lutolf M P,Hubbell J A. “Synthetic biomaterials as instructive extracellularmicroenvironments for morphogenesis in tissue engineering,” NatureBiotechnol. 2005; 23(1):47-55; Geckil H, Xu F, Zhang X H, Moon S,Demirci U. “Engineering hydrogels as extracellular matrix mimics,”Nanomedicine-UK. 2010; 5(3):469-484.]

In addition to including cells in the prepolymer layers, cells can beincluded in the nonpolymerizable fluids. In the experiment describedabove, endothelial cells were included in the interior wall of thetubule. Flow introduced subsequently through the hollow tube willencourage the formation of continuous monolayers on the inner wall ofthe tubule. Similar results would be obtained with renal epithelialcells in a matrix, mimicking the kidney basement membrane. Cells couldalso be included with the outer sheath layer that would attach to theouter surface of the tubule as it is formed to provide an even moredefined organization of multiple cell types.

Incorporation of the micro-blood vessels into tissue-on-chip systemsprovides for biomimetic delivery of nutrients and growth factors and forthe creation of biological gradients similar to those naturallyoccurring as different cell types are organized in complex tissues invivo. In one example, the micro-blood vessels are used to studyintramembranous ossification. Hierarchically-layered microfabricatedblood vessels are engineered through the combined use of hydrodynamicshaping and in situ polymerization. These micro-blood vessels withheterotypic cellular architecture are fabricated with diameters rangingfrom 5 μm to 1 mm and mean wall thickness between 0.5 μm and 500 μm. Themicro-blood vessels are integrated into the microchip bioreactor that isschematically shown in FIG. 33. Mesenchymal stem cells in abiocompatible hydrogel are introduced into the chamber surrounding themicro-blood vessels. Unlike other microchip tissue models [Gibbons M C,Foley M A, O'Halloran-Cardinal K. “Thinking inside the box: Keepingtissue-engineered constructs in vitro for use as preclinical models.Tissue Engineering: Part B,”. 2012; 19(1):17], the human vascularizedbone model described here incorporates both a free standing vasculaturenetwork and three-dimensional cell scaffold. The incorporatedvasculature provides the opportunity to model the mechanical andmaterial transport properties that dictate endothelial-mesenchymal stemcell interactions during intramembranous ossification. Analysis of bothvasculogenesis and osteogenesis will provide a biologically relevanthuman vascularized tissue model to investigate bone development andhealing.

This technique can also provide a multi-scale approach tomusculoskeletal tissue engineering that combines nanoscale molecularalignment, microengineered tissues and biomimetic scaffolds. Engineeredtissues for musculoskeletal repair are often limited by a lack ofvascularity, inability to replicate the complex bone microarchitectureand insufficient scaffold mechanical strength. Based on recent findingsthat unique material compositions and multiple cell types can beincorporated in a single micron-sized fiber via the hydrodynamicfocusing process described here, concentric-layered hydrogel constructswith spatially varying biomaterial layers can be used for organizing amixed-cell population into a complex muscle fiber or osteon. Theconcentric geometry of both muscle fibers and osteon can be achieved.

Anatomically and functionally, fibrous muscle tissue consists ofspatially distinct regions which each contain a distinct resident celltype and series of matrix tissue. Each zone is characterized by uniqueextra-cellular matrix compositions, mechanical properties and cellularorganization. Osteon, the functional unit of compact bone, arecylindrical structures that are typically several millimeters long andaround 0.2 mm in diameter. Each osteon consists of concentric layers, orlamellae, of compact bone tissue that surround a central canal. Byconcentrically layering biomimetic materials, microengineered musclefiber and osteon-like units can be formed. This will result inmineralized vascular tissues with controlled architectures on a varietyof functional and physiologically relevant scales. The combined use ofmicrofluidic shaping and in situ polymerization can produce multi-layerfibers with diameters ranging from 10 μm to 1 mm containing differentcell types. The fibers can be encased in an additional cell-free layerless than 10 μm thick to replicate the membranous structures thatencapsulate a muscle fiber or osteon. Microfluidic channels similar tothose in FIGS. 30 and 32 include additional fluid transporting regionsto focus pre-tissue solutions into the respective zones of a multi-layerfiber. Size-control in this regime is necessary to reproduce individualmuscle fibers. Groups of fibers can be bundled into large musculartissue models which contain both sarcolemma and endomysium.

FIG. 35 illustrates a model tissue microchip vascularized with thefibers configured as micro-blood vessels; the chip can be utilized forboth biomechanical and physiological analyses. The integrated tissuereservoir incorporates a 3-dimensional hydrogel matrix to supportvasculogenesis and 3-dimensional culturing of osteogenic cells. Such atissue model can incorporate micro-vessels or ducts or nerves embeddedin it as described above, and be used with 3-dimensional cultures of awide variety of tissue types. Inlet ports and outlet ports can directfluid through a hollow fiber, and/or provide nutrients either throughthe fiber itself or to a surrounding hydrogel matrix.

Each document mentioned herein is incorporated by reference.Furthermore, one of ordinary skill in the art will understand that manymodifications and variations of the present invention are possible inlight of the above teachings. It is therefore to be understood that,within the scope of the appended claims, the invention may be practicedotherwise than as specifically described.

What is claimed is:
 1. A model tissue comprising: a fiber comprising oneor more layers of polymer surrounding a hollow central lumen with aninner polymer surface nearest the lumen; a hydrogel matrix serving asthree-dimensional scaffold and at least partially surrounding the fiber;an inlet port and an outlet port at opposing ends of the fiber andoperable to direct fluid through the lumen; and living animal cellsdisposed in the model tissue and comprising at least endothelial cellsadhered to the inner polymer surface and in intimate contact with thelumen, and optionally a second cell type within the one or more layersof polymer and/or the three-dimensional scaffold, wherein the fiberremains intact during a condition of fluid flow through the lumensufficient for the fiber to serve as physiological model of a bloodvessel, and the fiber has an outer diameter of between 5 and 8000microns and wherein each individual layer of polymer has a thickness ofbetween 0.1 and 250 microns.
 2. The model tissue of claim 1, whereinsaid fiber is configured as a blood vessel, tissue duct, or nerve. 3.The model tissue of claim 1, comprising a plurality of said fibers,wherein a single inlet port and a single outlet port are operablyconnected to the plurality of fibers.
 4. The model tissue of claim 1,further comprising cells adhered to an exterior surface of said fiber.5. The model tissue of claim 1, wherein the fiber is free of cellattachment along a surface.
 6. The model tissue of claim 1, wherein thepolymer in at least one of said one or more layers is biodegradable. 7.The model tissue of claim 1, wherein the polymer in at least one of saidone or more layers comprises a material is selected from the groupconsisting of collagen, agarose, polyelectrolytes, chitosan, gelatin,polyethylene glycol, peptides, and combinations thereof.
 8. The modeltissue of claim 1, wherein at least one of said one or more layersfurther comprises a nucleic acid and/or a factor to modify cell growth,adhesion, and/or differentiation.
 9. A model tissue comprising: a fibercomprising: at least two concentric layers of polymer surrounding acentral lumen, and living animal cells disposed within at least one ofthe at least two concentric layers of polymer and comprising at leastendothelial cells adhered to the inner polymer surface and in intimatecontact with the lumen, and optionally within the lumen, wherein atleast two different animal cells are disposed in different layers ofpolymer and/or lumen, and an inlet port and an outlet port at opposingends of the fiber and operable to direct fluid through the lumen;wherein the fiber remains intact during a condition of fluid flowthrough the lumen sufficient for the fiber to serve as physiologicalmodel of a blood vessel, and the fiber has an outer diameter of between5 and 8000 microns and wherein each individual layer of polymer has athickness of between 0.1 and 250 microns, and wherein the fiber is atleast partially surrounded by a hydrogel matrix serving as athree-dimensional scaffold for cells, the hydrogel matrix being distinctfrom that of an outer-most of the at least two concentric layers ofpolymer.
 10. The model tissue of claim 9, wherein said fiber isconfigured as a blood vessel, tissue duct, or nerve.
 11. The modeltissue of claim 9, comprising a plurality of said fibers, wherein asingle inlet port and a single outlet port are operably connected to theplurality of fibers.
 12. The model tissue of claim 9, further comprisingcells adhered to an exterior surface of said fiber.
 13. The model tissueof claim 9, wherein the polymer in at least one of said one or morelayers is biodegradable.
 14. The model tissue of claim 9, wherein thepolymer in at least one of said one or more layers comprises a materialis selected from the group consisting of collagen, agarose,polyelectrolytes, chitosan, gelatin, polyethylene glycol, peptides, andcombinations thereof.
 15. The model tissue of claim 9, wherein at leastone of said one or more layers further comprises a nucleic acid and/or afactor to modify cell growth, adhesion, and/or differentiation.
 16. Themodel tissue of claim 1, wherein the living animal cells include humanumbilical vein endothelial cells surrounding the central lumen, primarycoronary artery smooth muscle cells in one of the layers, and humandermal fibroblasts in another one of the layers, exterior to the layerincluding the primary coronary artery smooth muscle cells.
 17. The modeltissue of claim 9, where said at least two different animal cellsinclude human umbilical vein endothelial cells surrounding the centrallumen, primary coronary artery smooth muscle cells in one of the layers,and human dermal fibroblasts in another one of the layers, exterior tothe layer including the primary coronary artery smooth muscle cells. 18.The model tissue of claim 1, further comprising a microfluidic chiphaving a chamber space defined by chamber walls housing the hydrogelmatrix, wherein the fiber passes through the chamber space and throughthe chamber walls with the lumen open at each end outside the chamberspace.
 19. The model tissue of claim 9, further comprising amicrofluidic chip having a chamber space defined by chamber wallshousing the hydrogel matrix, wherein the fiber passes through thechamber space and through the chamber walls with the lumen open at eachend outside the chamber space.